This invention relates to scintillator crystals and more particularly to scintillation crystal arrays, and a novel method for improving the detection, count efficiency and analysis of penetrating radioactive emissions. Because of the resulting potential improvements in spatial, spectral and temporal resolutions, the concept may be particularly useful in Positron Emission Tomography (PET).
Scintillation crystals are commonly used in non-invasive medical diagnostic techniques which utilize radiation emitting materials. These crystals are noted for their ability to emit pulses of visible light when ionizing radiation, such as gamma radiation, passes therethrough and interacts with atomic nuclei in the crystal. The pulses of emitted light (photons) are then detected by a photodetector device such as a photomultiplier tube (PMT) or a semiconductor photodiode (SPD). The effectiveness of the detector in diagnostic procedures depends on the ability to see and quantify the crystal light flashes with high spatial, spectral, and temporal precision. This in turn is dependent on brightness and rapidity of the generated flash which are functions of the type and geometry of the scintillation crystal.
The state-of-the-art PET gamma-ray detector used in commercial nuclear medicine cameras utilize a two-dimensional, discrete or pseudo-discrete array of long, narrow scintillation crystals which are coupled at a small end to PMTs with the opposite small end directed toward the gamma-ray source. The crystals are preferably long for high gamma-ray stopping power and narrow for high spatial resolution. An intermediate optical coupling medium is necessary in these designs at the scintillation crystal/PMT interface. Crystal surfaces in these designs are treated and coated with reflectors to preferentially direct light through internal reflections into the PMT located at one small end of the scintillation crystal. These prior art cameras typically quantify the scintillation light which reaches the small end of the scintillation crystal.
U.S. Pat. No. 5,091,650 describes a typical PMT/scintillator array arrangement. Since individual PMTs are expensive and rather bulky, the unit is usually constructed such that the number of scintillation crystals in a PET detector array is much larger than the number of PMTs used to read the light emissions. This scheme is termed "multiplexing." Typically, only four PMTs are required for one detector array unit, and an appropriate weighted mean of the 4 PMT signals determines the position of an event. Due to this multiplexing, there is a limit to the number of crystals in an array that can accurately be decoded by a given number of PMTs. This limits how narrow the crystals, or equivalently, how great the position resolution can be. For discrete crystal designs, one PMT per crystal is preferred, which is difficult for an array of narrow crystals because of size limitations of the PMT.
Another limitation in the standard PET detector design is that significant losses of scintillation light occur due to photon interactions with the crystal surfaces or reflective coatings on those surfaces. The results is that only a fraction of the scintillation light produced in the crystal reaches the photodetector. This light loss problem associated with standard photodetector readout at the end of the scintillating crystal worsens as the crystal is made narrower and longer or has unpolished side surfaces. This light loss problem together with the low quantum efficiency of the PMT photocathode for detecting the scintillation light produced limits the count efficiency and signal-to-noise ratio of both the crystal decoding scheme used to position and time a gamma-ray event, and the energy (spectral) resolution required to reduce gamma-ray scatter. Good scatter reduction is an important factor for improving image contrast between true structures of interest and the background present in the resulting PET images.
An additional related problem associated with the conventional end readout is that the light collection efficiency depends on the location within the crystal where light was created and thus, where the radiation interacted. This factor degrades the energy resolution. Also, there is roughly a 10-15% light loss at the interface between the crystal and PMT due to index of refraction mismatches, further degrading the signal to noise ratio.
As a result, commercially available PET detector arrays are bulky because of the PMT and inefficient because a significant portion of scintillation light generated never reaches the photodetector. This worsens the signal-to-noise ratio and count efficiency of conventional PMT-based PET detector designs and constrains the ultimate spatial, energy, and timing resolution available. Because of the use of PMTs, these designs are also very costly.